1. Field of the Invention
The present invention pertains to the field of detectors including, more specifically, to detectors capable of use in imaging systems.
2. Background
In a typical known embodiment of a reverse-geometry x-ray imaging system, an output signal from a detector is applied to the z-axis (luminance) input of a video monitor. This signal modulates the brightness of the viewing screen. The x and y inputs to the video monitor are typically derived from the signal that effects deflection of the electron beam of an x-ray tube. Therefore, the luminance of a point on the viewing screen is inversely proportional to the absorption of x-rays passing from the source, through the object, to the detector.
Time and area distributions of x-ray flux follow a Poisson distribution and have an associated randomness that is unavoidable. The randomness is typically expressed as the standard deviation of the mean flux, and equals its square root. The signal-to-noise ratio of a x-ray image under these conditions is equal to the mean flux divided by the square root of the mean flux. i.e., for a mean flux of 100 photons, the noise is +/-10 photons, and the signal-to-noise ratio is 10.
In many reverse-geometry scanning beam x-ray systems, the spatial resolution of the resulting image is in large part determined by the capture area of one of the detector elements. Generally speaking, a non-segmented detector with a small capture area can provide high spatial resolution and poor collection efficiency (i.e., the ratio of the meaningful photons passing through the object to the total number of photons passing through the object), while a non-segmented detector with a large capture area provides high collection efficiency and poor spatial resolution.
A medical imaging system should provide low patient dosage, high spatial resolution and a refresh rate of up to about 30 times per second which is the refresh rate of a standard video display--all at the same time. The spatial resolution and the signal-to-noise ratio of x-ray images formed by known reverse-geometry x-ray imaging systems are dependent, to a large extent, upon the size of the sensitive area of the detector element. If the sensitive area of the detector element in these systems is increased, more of the diverging rays are detected, effectively increasing sensitivity and improving the signal-to-noise ratio. At the same time, however, in some medical applications, larger detector areas tend to reduce the attainable spatial resolution as the "pixel" size (measured at the plane of the object to be imaged) becomes larger. This is so because of the distance typical objects to be imaged in some medical applications (e.g., structures internal to the human body) are from the x-ray source. These issues may be addressed to a certain extent by increasing the x-ray photon flux.
In the medical field, several conflicting factors, among them patient dosage, frame rate (the number of times per second that the image refreshed), and resolution of the image of the object, often add additional constraints. For example, a high x-ray flux may provide high resolution and a high frame rate, yet result in an unacceptably high x-ray dosage to the patient and attending staff.
While on the other hand, lower dosages may be achieved from known systems at the cost of a resolution image or a lower refresh rate. In known medical x-ray imaging systems, therefore, the detector element area typically selected to effect a compromise between resolution and sensitivity given the other constraints.
Solid state x-ray detectors for x-ray imaging systems are known. An example of a solid state x-ray detector is disclosed in U.S. Pat. 5,379,336 to Kramer et al. Kramer et al. discloses a hybrid detector array with many thousands of individual detector pixels on a single semiconductor substrate interconnected to a corresponding readout with individual amplifiers and signal condition circuits for each pixel and multiplexor output. The detector pixels of the type disclosed in Kramer et al. generate electrical signals in response to x-ray photons. The detector pixels disclosed in Kramer et al. are 30 .mu.m by 30 .mu.m. These detector pixels are therefore of a small size to increase spatial resolution. One problem associated with having detector pixels of a small size is that more than one detector pixel in a given area may be effected by a single x-ray photon. This situation can affect the accuracy of an image of an object under investigation.
There is a need for a x-ray detector that has high resolution and collection efficiency and that provides information in a way that is easy to process in an efficient manner. There is additionally a need for a x-ray detector that minimizes the possibility of signals indicating the false or phantom detection of x-ray radiation while minimizing the x-ray dosage to the object under investigation.